The present invention relates to an image sensor. It also relates to a method of fabricating such a sensor. The invention is of particular use in the field of radiological imaging where the image sensors include a scintillator for converting a probe radiation, for example an X- or Gamma-range radiation, into a normally visible detection radiation.
Already known in the field of image sensors are electronic X-radiation sensors including radiological image intensifying screens disposed facing a detector and receiving an X-radiation. The detectors most commonly used are, in the field of radiology, cameras with targets, strip assemblies or matrices of charge-coupled devices (CCD), or even matrices of CMOS (Complementary Metal Oxide Semiconductor) detectors. These detectors are relatively ineffective for directly detecting X rays, which is why they are, for this purpose, often associated with scintillators for converting the X-range radiations into radiations in the spectrum of sensitivity of the detector, for example in the visible spectrum.
The material used to obtain the conversion, that is, the material of the scintillator, is often gadolinium oxysulphide. This is used in the form of a thin film, typically of the order of 50 to 300 micrometers. This film is made of particles of this material joined by a binder. The emission of visible light through the entire thickness, and in all the directions, of this material results in a loss of resolving power of the detector and therefore of the X-radiation sensor.
Thallium-doped caesium iodide Csl, in needle form, offers an interesting alternative for a greater light efficiency associated with a waveguide effect of the needles, the typical dimensions of the sections of which range from 3 to 6 micrometers. Implementations are thus known in which the scintillator of the input screen is made of caesium iodide deposited by vacuum evaporation onto a substrate, the evaporation possibly taking place on a cold or hot substrate. The needles are oriented perpendicularly to the surface of the substrate bearing them. They are only partially adjoined to each other. They thus offer a porosity of 20 to 25%. These pores, or interstices, filled with air, associated with the favourable refraction index of the Csl (1.78), provide a channelling of the visible photons emitted in each needle and give a higher sensitivity and resolving power.
FIG. 1 diagrammatically represents a radiological image sensor comprising a scintillator 1, and a detector 2. The scintillator 1, comprises a substrate 3, for example made of carbon, covered with a layer of a luminescent material 4, for example caesium iodide, which takes the form of needles 5 (or asperities). The detector 2, for example a strip of charge-coupled devices, comprises elements 6 sensitive to the visible radiation. The thickness of the layer of luminescent material 4 is constant over all the surface of the substrate.
The substrate 3 receives a stream of X photons (probe radiation) symbolized by solid-line vertical arrows. Broken lines in the figure represent examples of paths 10, 11, followed in the needles 5 of caesium iodide, by the visible radiation corresponding to the incident X photons. The visible radiation (detection radiation) emerging from the luminescent material 4, illuminates the detector 2. The normal paths 10 have a direction roughly parallel to that of the X photons, and they emerge from the scintillator via the ends of the needles 5 made of caesium iodide. There is also lateral diffusion of the visible radiation carried in the needles 5 made of caesium iodide, as is indicated in the figure by the reference 11.
The light efficiency of the scintillator is defined as the ratio between the energy of the incident probe radiation on the scintillator and the energy of the detection radiation emitted by the scintillator.
The spatial resolution of the radiological image sensor depends on the capacity of the needles 5 of caesium iodide to effectively channel the visible radiation. This capacity is a function of the thickness of the layer of luminescent material 4. An increase in the thickness of the layer of luminescent material 4 results in a deterioration of the spatial resolution of the radiological image sensor.
Moreover, when the sensitive elements 6 of the detector 2 are illuminated by a visible radiation, the sensitive elements 6 create electrons which they store temporarily within themselves, in order for them to be converted into an electrical image outside the detector. The sensitive elements 6 have an electron storage capacity which can, in certain cases, be less than the number of electrons generated by a visible radiation emitted locally by the scintillator 1. To avoid local saturations of stored electrons resulting in saturations on the electrical image, it may be necessary either to modify the spectrum of the visible radiation arriving on the detector 2, or even to attenuate the energy of the visible radiation emitted by the luminescent material 4, which is tantamount to reducing the light efficiency of the scintillator.